1. Field of the Invention
The present invention relates in general to a gradient system for use in magnetic resonance tomography (MRT) for examining patients. The present invention relates, in particular, to a switchable gradient system and to a method for calculating such a system, in the case of which so-called booster coils are used.
2. Description of the Prior Art
MRT is based on the physical phenomenon of nuclear spin resonance and has been used successfully as an imaging method for over 15 years in medicine and in biophysics. In this method of examination, the object is exposed to a strong, constant magnetic field. This aligns the nuclear spins of the atoms in the object, which were previously oriented randomly. Radio-frequency energy can excite these xe2x80x9corderedxe2x80x9d nuclear spins to a specific oscillation (resonant frequency). In MRT, this oscillation generates the actual measurement signal (RF response signal), which is picked up by suitable receiving coils.
Having exact information relating to the respective origination location of the RF response signal (location information or location coding) is a precondition for the image reconstruction. This location information is obtained by means of additional magnetic fields (magnetic gradient fields) relative to the static magnetic field along the three spatial directions. These gradient fields are small by comparison with the main field and are generated by additional resistance coils in the patient opening of the magnet. Due to the gradient fields, the overall magnetic field differs in each volumetric element and therefore so is the resonant frequency. If a defined resonant frequency is irradiated, it is therefore possible to excite only the atomic nuclei that are situated at a location at which the magnetic field fulfills the resonance condition. A suitable change in the gradient fields allows the location of such a volumetric element where the resonance condition is fulfilled to be displaced in a defined fashion, and thus for the desired region to be scanned. The gradient fields are therefore switched repeatedly in an MR sequence for excitation (coding) and reading out (detection) of the nuclear resonance signals.
This allows a free choice of the layer to be imaged, as a result of which it is possible to obtain tomographic images of the human body in all directions.
The basic design of an MRT apparatus is illustrated in FIG. 4. A basic field magnet 23 (for example an axial superconducting air-coil magnet with active stray field screening) which generates a homogeneous magnetic basic field in an interior space. The superconducting magnet 23 contains superconducting coils which are located in liquid helium. The basic field magnet 23 is surrounded by a double-shell tank which is made of stainless steel, as a rule. The inner tank, which contains the liquid helium and serves in part as winding body for the magnet coils, is suspended at the outer tank, which is at room temperature, via fiber-glass-reinforced plastic rods which are poor conductors of heat. A vacuum prevails between inner and outer tanks.
The cylindrical gradient coil 25 in the inner space of the basic field magnet 23 is inserted concentrically into the interior of a support tube by means of support elements 24. The support tube is delimited externally by an outer shell 26, and internally by an inner shell 27.
The gradient coil 25 has three windings which generate respective gradient fields, each field being proportional to the current in each coil and being orthogonal to one another. As illustrated in FIG. 5, the gradient coil 25 has an x-gradient coil 28, a y-gradient coil 29 and a z-gradient coil 30, which are respectively wound around the coil core 31 and thus generate a gradient field, respectively in the directions of the Cartesian coordinates x, y and z. Each of these coils is fitted with a dedicated power supply unit in order to generate independent current pulses with accurate amplitudes and timing in accordance with the sequence programmed in the pulse sequence controller 20. The required currents are at approximately 250 A. Since the gradient switching times are to be as short as possible, current rise rates and current fall rates (slew rates) on the order of magnitude of 250 kA/s are necessary. A high gradient power is required in order to obtain images with high spatial resolution and a short measuring time. According to the current state of the art, the gradient intensities are 30-60 mT/m for switching times of 100-500 xcexcs.
However, these high rates of change in the magnetic field in the body of the patient, at times, can cause painful peripheral nerve stimulations. The threshold for peripheral nerve stimulations scales with the magnitude of the homogeneity volume (DSV=Diameter of Spherical Volume, or field mode), which is fixed by the gradient system.
For this reason, and in order to do justice to different applications in MRTxe2x80x94in particular in functional imagingxe2x80x94it is necessary to lend the MRT machine DSV flexibility. The technically possible system power, in conjunction with avoidance of peripheral nerve stimulations, can be fully exploited according to the current state of the art by using a gradient system that has a number of field characteristics (field modes). The field characteristic, or the field mode, describes a generally spherical region in the interior of the homogeneous basic field in which the gradient deviates by less than 5% from the reference value at the coil center. The radius and quality of the corresponding homogeneity region definitively determine the essential system properties of the gradient system such as switching time, maximum gradient intensity and stimulation threshold. According to the current state of the art, they can be changed in discrete steps by using switchable gradient coils.
According to the current state of the art, a switchable gradient system with a number of field characteristics can be implemented in various ways respectively having different advantages and disadvantages:
A) by combining or integrating a number of (completely shielded) coil sections.
B) Modular conductor bundling, by combining suitable conductor bundles within a coil plane for a discrete number of field characteristics.
In order, for example, to generate two different field characteristics, in method A two different (actively shielded) coils are interleaved. Different field modes can be obtained by appropriate electric interconnection of the two coils.
FIG. 6a schematically illustrates the idealized z-direction field pattern of two whole-body gradient fields DSV1, DSV2 with different homogeneity radii. The two fields are generated by a gradient system of a whole-body tomography apparatus with two independent coil sections in the z-direction. FIG. 6b shows a transverse section (x-y plane) through the whole-body tomography apparatus.
The gradient system is shielded from the outside and from the superconducting basic field magnet by a cryoshield 32 (referred to as the tank above). The system uses a large whole-body coil 33, which produces a correspondingly large spherical homogeneity volume (DSV1) 34.
Disadvantages of this large-volume whole-body coil 33 are a high inductance and a high stimulation effect. These two disadvantages can be compensated by the use of a second smaller coilxe2x80x94the so-called insert coil 35. By energizing the insert coil 35, a small, elliptical homogeneity volume (DSV2) 36 inside the large homogeneity volume of the basic field magnet is typically obtained.
As can be seen from FIG. 6b, each of the two coil sections 33, 35 occupies a hollow cylinder of a certain thickness. The radii of the respective hollow cylinders are different as a rule, and the coil sections are therefore located on different winding planes. As already indicated and as can be seen in FIG. 4bxe2x80x94this leads to a reduction in the inside diameter of the gradient tube and to a reduction in the patient space.
Such a design is therefore compatible with the current MRT machine standards only in a restricted fashion. The radius of the insert coil may therefore not be selected to be excessively small.
Method A) therefore has the disadvantage that it can only be used, given the currently customary current strengths ( greater than 400A) and load cycles, if the inside diameter of the gradient tube may be reducedxe2x80x94owing to the greater space requirement for additional coil sections and for additional cooling planes. However, since a defined magnet diameter and thus a defined space requirement for the MRT machine components have gained international acceptance, changing specific components would be attended by a substantial extra outlay and costs connected therewith in the case of current MRT systems.
In method B), conductor bundles are combined within a coil plane and interconnected in different ways such that desired field characteristics can be generated.
The difference between the modular conductor bundling and method A) is that the conductor bundling relates to the winding plane of a coil. No independent coils are used, but existing coils are suitably divided.
In this case, the field property of each individual conductor (turn) is firstly considered in computational terms. Jointly usable conductor regions (conductor bundles) can be identified by subsequently comparing different current density distributions that are generated computationally by different combinations of individual conductors. The selection of suitable conductor bundles is performed in each case for a desired characteristic homogeneity volume. Methods for identifying optimum conductor combinations are, for example, Simulated Annealing or Genetic Algorithms.
The principle of conductor bundling in the case of transverse gradient coils (saddle coils) or longitudinal gradient coils (Maxwell coils) is illustrated schematically in FIGS. 7a and 7b, respectively.
In order to calculate the corresponding coil, a known iterative optimization method (target field) is used. In this method, the desired magnetic field pattern (the target field) on a desired DSV geometry (cylinder surface, ellipsoid: 37, 38, 39, etc.) is prescribed in the coil interior, and the current density distribution on the coil surface is derived therefrom. The contour lines of the integral of the calculated current density are used to generate the conductor track layout.
It may be seen from FIG. 7a) that the conductor tracks of a transverse coil octant form semi-ellipses whose number rises in general with a rising field homogeneity requirement (39xe2x86x9238xe2x86x9237) in the axial direction.
In a first step, the conductors are packed in the longitudinal (axial) direction. As can be seen in FIG. 7b, the longitudinal extent of the homogeneity ellipsoid scales directly with the longitudinal extent of the conductor packet (37xe2x86x9240; 38xe2x86x9241; 39xe2x86x9242).
In general, the resulting field characteristics have inadequate rotational field shielding and inadequate linearity. Consequently, the packing is refined in a second step.
In the second step, individual conductors 43 are specifically permutated over winding packets. The quality of this correction is a function of the number of the available conductor loops.
However, with increasing shortening of the homogeneity ellipsoid (39xe2x86x9238xe2x86x9237) there is also a shortening in the coil generating the homogeneity ellipsoid, and thus a decrease in its gradient intensity: the field efficiency is reduced (the value that guarantees a defined field strength given a defined energization of the coil). The field efficiency denotes the gradient intensity that is obtained when a defined current is impressed on a coil (for example 1Axe2x86x9280 xcexcT/Am).
Thus, method B) has the disadvantage that the conductor bundling entails an extremely complex interconnection; and the field efficiency decreases with decreasing imaging volume (field mode with small homogeneity volume).
It is an object of the present invention to provide a switchable gradient system and a method for constructing such a system, in which the above disadvantages become irrelevant or are eliminated, that is, in which the conductor bundling for the different field modes is simplified to facilitate production and the field efficiency is compensated in the case of small field modes.
This object is achieved in accordance with the invention in a magnetic resonance tomography apparatus having a gradient coil system which includes at least one gradient coil arrangement for generating a gradient field in a spatial direction. This gradient coil arrangement includes at least one gradient coil in the form of a primary coil with different conductor bundles, and a secondary coil for shielding the primary coil, as well as at least one booster coil that can likewise be constructed from different conductor bundles. The primary coil, secondary coil and booster coil are interconnected so that different homogeneity volumes can be produced at the center of the gradient coil system.
In order to produce a desired homogeneity volume with enhanced performance characteristics, according to the invention one or more modules of the primary coil are connected in series to the secondary coil and one or more booster coils.
In accordance with the invention, each booster coil is arranged on a winding plane between the primary coil and secondary coil.
It can be advantageous in this case for the booster coil to be of modular construction.
The above object also is achieved in accordance with the invention in a method for constructing the gradient system of a magnetic resonance tomography machine including the following steps:
i) calculating conductor tracks for the gradient coil for a prescribed homogeneity volume by using an optimization method, the design prescription of the primary coil and the secondary coil thereby being obtained, and
ii) determining a subset of conductor tracks of the primary coil by means of which another prescribed homogeneity volume is produced together with the secondary coil and at least one booster coil likewise calculated by the above optimization method.
It can be advantageous in this case as well for the booster coil to be of modular construction.
According to the invention, an interconnection of the secondary coil with conductor bundles of the booster coil and with conductor bundles of the primary coil can be calculated so that different homogeneity volumes can be produced.
In the above method, the winding plane of the booster coil is arranged between the winding plane of the secondary coil and the winding plane of the primary coil.